Magnetic resonance (MR) imaging of internal body tissues may be used for numerous medical procedures, including diagnosis and surgery. Generally, an MRI system 100, as depicted in FIG. 1, includes a static-field magnet 102, one or more gradient-field coils 104, a radio-frequency (RF) transmitter 106, and an RF receiver (not shown). (In some embodiments, the same device is used alternately as RF transmitter or receiver.) The magnet includes a region 108 for receiving a patient 110 therein, and provides a static, relatively homogeneous magnetic field over the patient, which causes hydrogen nuclei spins to align with and precess about the general direction of the magnetic field. The spin alignment creates a net magnetization in the tissue that depends, generally, on the type of tissue and can, thus, be used to create contrast in an MR image. Time-variable magnetic field gradients generated by the gradient-field coils 104 are superposed with the static magnetic field so as to encode spatial information by spatio-temporally varying the precession frequency of the spins. The RF transmitter 106 transmits RF pulse sequences over the patient 110 to cause some of the aligned spins to alternate between a temporary high-energy non-aligned state and the aligned state, thereby inducing an RF response signal called the MR echo or MR response signal. To obtain an MR image, the MR response signal is integrated over the entire (two- or three-dimensional) imaging region and sampled by the RF receiver to produce a time series of response signals that constitute the raw image data. This raw data is passed on to a computation unit 112. Each data point in the time series can be interpreted as the value of the Fourier transform of the position-dependent local magnetization at a particular point in k-space (i.e., wavevector space), where the wavevector k is a function of the time development of the gradient fields. Thus, by Fourier-transforming the time series of the response signal, the computation unit 112 can reconstruct a real-space image of the tissue (i.e., an image showing the measured magnetization-affecting tissue properties as a function of spatial coordinates) from the raw data. The real-space MR image may then be displayed to the user.
The MRI system 100 may be used to plan a medical procedure as well as to assist in locating and guiding medical instruments and monitor treatment progress during the procedure. For example, a medical procedure can be performed on a patient using a medical instrument while the patient is in the MRI machine. The medical instrument may be inserted into the patient, or used non-invasively, i.e., placed externally to the patient while creating a therapeutic or diagnostic effect in the tissue. MRI may be used to image an anatomical region of the patient, locate a treatment target within the region, monitor the location of the medical instrument (or the focus of its effects) relative to the target (preferably in real time), and/or monitor the temperature in and surrounding the target tissue.
For instance, the medical instrument can be a focused ultrasound device 114 that is located outside a patient's body and focuses ultrasonic energy into the patient's body. Ultrasound penetrates well through soft tissues and, due to its short wavelengths, can be focused to spots with dimensions of a few millimeters; therefore, it can be used for highly localized non-invasive surgery—for example, to ablate, coagulate, or otherwise necrose cancerous tissue without causing significant damage to surrounding healthy tissue. An ultrasound focusing system generally utilizes an acoustic transducer surface, or an array of transducer surfaces, to generate an ultrasound beam. The transducer may be geometrically shaped and positioned such that the ultrasonic energy is focused at a “focal zone” corresponding to the target tissue mass within the patient. During wave propagation through the tissue, a portion of the ultrasound energy is absorbed, leading to increased temperature and, eventually, to cellular necrosis—preferably at the target tissue mass in the focal zone. The individual surfaces, or “elements,” of the transducer array are typically individually controllable, i.e., their phases and/or amplitudes can be set independently of one another (e.g., using a “beamformer” with suitable delay and amplifier circuitry for the elements), allowing the beam to be steered in a desired direction, focused at a desired distance, and its beam profile to be conformed to a desired shape. Thus, the focal zone can be rapidly displaced and/or reshaped by independently adjusting the amplitudes and phases of the electrical signal input into the transducer elements; the transducer elements, in other words, are operable as a phased array.
During MR-guided focused-ultrasound (MRgFUS) treatment, patient motion (such as periodic motion due to respiration or random movements) can pose a considerable challenge to therapeutic efficacy and safety. Compensation for motion is necessary to ensure that the ultrasound beam remains focused on the target and does not damage the surrounding healthy tissues. In MRgFUS systems, motion compensation is generally accomplished by tracking the target in the images and steering the ultrasound beam based on the tracked position. One approach to target tracking involves directly determining the coordinates of the target, or of easier identifiable “anatomical landmarks” at fixed locations relative to the target, in the images. In an alternative approach, the relative shifts between successive images are determined by correlating one image with a large number of computationally shifted copies of the other image, and selecting the shifted image that provides the best match. In either case, significant image-processing time is expended to determine the target location. Thus, if such image processing is performed on the images acquired during treatment, the effective imaging rate is typically increased significantly, often impeding real-time motion compensation. This may cause beam-targeting inaccuracies and/or necessitate treatment interruption to correct for any misalignment due to displacement of the target tissue or organ.
To avoid these problems and facilitate target tracking in real time, a library of reference images covering different stages within the anticipated range of patient motion may be acquired and analyzed prior to treatment. The location of the target (or other object of interest) within each reference image is stored along or in association with the respective image, e.g., in an integrated reference record. As actual treatment proceeds, the images acquired in real time are correlated against the reference images in the library to determine matches based on image similarity. The location of the target region in the acquired treatment image is then inferred from the locational information associated with the corresponding reference image. Because image matching is, generally, computationally less involved then detecting and localizing objects within an image, this approach can achieve significant savings in processing time during treatment, thus facilitating real-time tracking.
Motion compensation is also relevant in MR-based thermometry (i.e., the generation of temperature maps of a monitored anatomical region from MR images thereof), where it can likewise benefit from a reference library acquired prior to treatment. Thermometry facilitates monitoring the progress of thermal treatment of target tissue, e.g., to ensure that non-target tissues are not inadvertently heated beyond clinically tolerable levels. Among various methods available for MR thermometry, the proton resonance frequency (PRF) shift method is often optimal due to its excellent linearity with respect to temperature change, near-independence from tissue type, and temperature map acquisition with high spatial and temporal resolution. The PRF shift method is based on the phenomenon that the MR resonance frequency of protons in water molecules changes linearly with temperature (with a constant of proportionality that, advantageously, is relatively constant between tissue types). Since the frequency change with temperature is small, only −0.01 ppm/° C. for bulk water and approximately −0.0096 to −0.013 ppm/° C. in tissue, the PRF shift is typically detected with a phase-sensitive imaging method in which the imaging is performed twice: first to acquire a baseline (or reference) PRF phase image prior to a temperature change and then to acquire a second phase image—i.e., a treatment image—after the temperature change, thereby capturing a small phase change that is proportional to the change in temperature. A map of temperature changes may then be computed from the (reconstructed, i.e., real-space) images by determining, on a pixel-by-pixel basis, phase differences between the baseline image and the treatment image, and converting the phase differences into temperature differences based on the PRF temperature dependence while taking into account imaging parameters such as the strength of the static magnetic field and echo time (TE) (e.g., of a gradient-recalled echo). Further, if the temperature distribution in the imaged area at the time of acquisition of the baseline image is known, the temperature-difference map can be added to that baseline temperature in order to obtain the absolute-temperature distribution corresponding to the treatment image.
The ability to obtain a temperature (difference) map for a treatment image depends on the existence of a suitable reference image, i.e., an image that, up to the temperature distribution, reflects the imaging conditions, including the location of the object(s) of interest, as they exist at the time the treatment image is acquired. If the region of interest is stationary (as are, typically, e.g., the prostate and uterine tract), a single reference (or baseline) image may suffice. Typically, however, the patient—and with him the region to be monitored or one or more organs therein—moves during treatment; such motion can be periodic (e.g., due to respiration) or sporadic and random. In this case, a library of reference images covering the range of motion may be acquired prior to treatment (e.g., heating), and an absolute-temperature map may optionally be stored along with each reference image (e.g., forming a reference record including the MR image and temperature map for each stage of motion). To obtain the proper reference (or baseline) image for a new treatment image, a correlation or other suitable image-selection technique is performed against the library to find the baseline image best aligned, spatially, with the treatment image. The selected baseline image and treatment image are processed as described above to determine the changes in temperature, and an absolute-temperature map for the treatment image is computed based on the temperature map corresponding to the baseline image and the image-to-image phase changes. This method is often referred to as multi-baseline thermometry. Additional algorithms (e.g., to account for phase wrapping, to correct for drift in the static magnetic field, or to integrate measurements from multiple resources such as multiple MR channels) may also be applied.
Although thermometry and object tracking for the purpose of beam steering involve very different operations and technical constraints, in both cases, the ability to accommodate movement may be critical, and may depend on the robustness of the reference library. When the patient's movement exceeds what is anticipated (i.e., the target region in the treatment image is no longer in a region covered by the reference library and/or a baseline image is not found because the image similarities between the treatment image and reference images are insufficient), it may be necessary to stop treatment in order to correct misalignment due to displacement of the target tissue or organ. Movements of anatomical structures other than the target can also negate the usefulness of the reference library by disturbing the electromagnetic field (which directly affects the phase map of the image) so that the resulting treatment images do not directly correspond in detail to the reference images. Unless movement and other changes exceeding the coverage of the reference library are transient and brief, treatment must typically be halted to allow for recalibration and realignment of the treatment device and for the acquisition and/or processing of new reference images. The result is inconvenience and delay.
Accordingly, there is a need for the efficient extension of the reference library when reference images do not match newly obtained images (due to movements, changes in imaging parameters, or other factors), preferably so that treatment can be continuously performed without interruption.